Pressure-based system and method for determining cardiac stroke volume

ABSTRACT

Cardiac stroke volume (SV) of a subject is estimated as a function of a value derived from a measured arterial pressure waveform. The value may be the standard deviation, or a function of the difference between maximum and minimum pressure values, or a function of either the maximum value of the first time derivative or the absolute value of the minimum of the first time derivative of the pressure waveform, or both, or a function of the magnitude of one or more spectral components of the pressure waveform at a frequency corresponding to the heart rate. Cardiac output is then estimated as the product of the subject&#39;s heart rate and SV, scaled by a calibration constant. Arterial pressure may be measured invasively or non-invasively.

RELATED APPLICATIONS

The present application is a divisional of U.S. Ser. No. 10/728,705filed on Dec. 5, 2003 now U.S. Pat. No. 7,220,230 entitledPressure-Based System and Method for Determining Cardiac Stroke Volumewhich is incorporated herein by reference in its entirety.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to a method for determining the stroke volume(SV) and, hence, any other SV-related value such as cardiac output of ahuman or animal subject, as well as to a system that implements themethod.

2. Background Art

Cardiac output (CO) is an important indicator not only for diagnosis ofdisease, but also for “real-time” monitoring of the condition of bothhuman and animal subjects, including patients. Few hospitals aretherefore without some form of conventional equipment to monitor cardiacoutput.

One basis for almost all common CO-measurement systems is the well-knownformula CO=HR·SV, where SV is the stroke volume and HR is the heartrate. SV is usually measured in liters and HR is usually measured inbeats per minute. This formula simply expresses that the amount of bloodthe heart pumps out in a minute is equal to the amount it pumps out onevery beat (stroke) times the number of beats (strokes) per minute.

Since HR is easy to measure using any of a wide variety of instruments,the calculation of CO usually depends on some technique for estimatingSV. Many suitable techniques—both invasive and non-invasive, as well asthose that combine both—are in use and even more have been proposed inthe literature.

One invasive way to determine cardiac output (or, equivalently, SV) isto mount some flow-measuring device on a catheter, and then to place thecatheter into the subject and to maneuver it so that the device is in ornear the subject's heart. Some of such devices inject either a bolus ofmaterial or energy (usually heat) at an upstream position, such as inthe right atrium, and determine flow based on the characteristics of theinjected material or energy at a downstream position, such as in thepulmonary artery. Patents that disclose implementations of such invasivetechniques (in particular, thermodilution) include:

U.S. Pat. No. 4,236,527 (Newbower et al., 2 Dec. 1980);

U.S. Pat. No. 4,507,974 (Yelderman, 2 Apr. 1985);

U.S. Pat. No. 5,146,414 (McKown, et al., 8 Sep. 1992); and

U.S. Pat. No. 5,687,733 (McKown, et al., 18 Nov. 1997).

Still other invasive devices are based on the known Fick technique,according to which CO is calculated as a function of oxygenation ofarterial and mixed venous blood. In most cases, oxygenation is sensedusing right-heart catheterization. There have, however, also beenproposals for systems that measure arterial and venous oxygenationnon-invasively, in particular, using multiple wavelengths of light, butto date they have not been accurate enough to allow for satisfactory COmeasurement on actual patients.

Invasive techniques have some disadvantages, the main one of which is ofcourse that catheterization of the heart is more dramatic to thepatient, especially considering that the subjects (especially intensivecare patients) on which it is performed are often already in thehospital because of some actually or potentially serious condition.Invasive methods also have less obvious disadvantages: Some techniquessuch as thermodilution rely on assumptions, such as uniform dispersionof the injected heat, that affect the accuracy of the measurementsdepending on how well they are fulfilled. Moreover, the veryintroduction of an instrument into the blood flow may affect the value(for example, flow rate) that the instrument measures.

There has therefore been a long-standing need for some way ofdetermining CO that is both non-invasive—or at least as minimallyinvasive as possible—and accurate. One blood characteristic that hasproven particularly promising for accurately determining COnon-invasively is blood pressure.

Most known blood-pressure-based systems rely on the so-called pulsecontour method (PCM), which calculates as estimate of GCO fromcharacteristics of the beat-to-beat pressure waveform. In the PCM,“Windkessel” (German for “air chamber”) parameters (characteristicimpedance of the aorta, compliance, and total peripheral resistance) areused to construct a linear or non-linear, hemodynamic model of theaorta. In essence, blood flow is analogized to a flow of electricalcurrent in a circuit in which an impedance is in series with aparallel-connected resistance and capacitance (compliance). The threerequired parameters of the model are usually determined eitherempirically, through a complex calibration process, or from compiled“anthropometric” data, that is, data about the age, sex, height, weight,etc., of other patients or test subjects. U.S. Pat. No. 5,400,793(Wesseling, 28 Mar. 1995) and U.S. Pat. No. 5,535,753 (Petrucelli, etal., 16 Jul. 1996) are representative of systems that rely on aWindkessel circuit model to determine CO.

PCM-based systems can monitor CO more or less continuously, with no needfor a catheter to be left in the patient. Indeed, some PCM systemsoperate using blood pressure measurements taken using a finger cuff. Onedrawback of PCM, however, is that it is no more accurate than the rathersimple, three-parameter model from which it is derived; in general, amodel of a much higher order would be needed to faithfully account forother phenomena, such as the complex pattern of pressure wavereflections due to multiple impedance mismatches caused by, for example,arterial branching. Because the accuracy of the basic model is usuallynot good enough, many improvements have been proposed, with varyingdegrees of complexity.

The “Method and apparatus for measuring cardiac output” disclosed bySalvatore Romano in U.S. Published Patent Application 20020022785 A1represents a different attempt to improve upon PCM techniques byestimating SV, either invasively or non-invasively, as a function of theratio between the area under the entire pressure curve and a linearcombination of various components of impedance. In attempting to accountfor pressure reflections, the Romano system relies not only on accurateestimates of inherently noisy derivatives of the pressure function, butalso on a series of empirically determined, numerical adjustments to amean pressure value.

What is needed is a system and method of operation for estimating COthat is robust, simple, and accurate and that does not requireanthropometric values or repeated calibrations. This invention meetsthis need.

SUMMARY OF THE INVENTION

A parameter proportional to the cardiac stroke volume (SV) of a patientis determined by sensing an input signal that either directly indicatesor is proportional to arterial blood pressure. The sensor used to sensethe input signal may be either invasive or non-invasive. The standarddeviation of the input signal is then calculated over a measurementinterval and an estimate of SV is then calculated as a function of thestandard deviation of the input signal. SV may be also computed as theproduct of the standard deviation and a calibration factor. In anexemplifying processing system that implements the method, one or morecomputer-executable software modules are included for carrying out thevarious calculations.

Any cardiac value derived from SV may also use the invention todetermine an SV estimate to be used for calculating the value. Forexample, the method according to the invention may be used to calculatean estimate of cardiac output (CO). In such an application of theinvention, any known mechanism (for example, a hardware monitor and/orsoftware algorithm) is used to measure the patient's heart rate (HR).The current cardiac output of the patient is then estimated, forexample, by calculating the product of HR and the standard deviation andscaling the product by a calibration constant.

In CO applications of the invention, the calibration constant may bedetermined using different techniques, both invasive and non-invasive.To calculate the calibration constant, a calibration cardiac outputvalue is measured and the calibration constant is approved as thequotient between a calibration cardiac output estimate and the productof the heart rate and the standard deviation.

The measurement interval may extend over more than one cardiac cycle,for example, to cover a time window that is multiple cardiac cycleswide. A single standard deviation value of the input signal may becalculated over the whole interval, or component standard deviationvalues may be calculated and then averaged (using the mean, median,etc.) for each of a plurality of sub-intervals to form a final compositestandard deviation value that can be used in calculating the estimate ofthe cardiac stroke volume.

Various optimizations may be included in different embodiments of theinvention. For example, for each of a plurality of cardiac cycles, amean pressure value can be calculated and the measurement interval canthen be adjusted as a function of change in the mean pressure value.

If needed, for example, to remove the effect of potential drift in meanpressure over the measurement interval(s), the input signal may behigh-pass filtered before standard deviation is calculated.

Standard deviation may be calculated in different ways, having differentdegrees of statistical accuracy. For example, the input signal may bediscretized over the measurement interval, and then a standard algorithmmay be applied to determine an estimate of standard deviation from thesample values. As an alternative, standard deviation may be approximatedas a function of the difference between the maximum and minimum values.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is an illustrative example of a complex blood pressure curve overone beat-to-beat heart cycle

FIG. 2 is a block diagram showing the main components of a systemaccording to the invention.

DETAILED DESCRIPTION Introduction

In broadest terms, the invention involves the determination of a cardiacvalue such as stroke volume (SV), and/or a value derived from SV such ascardiac output (CO), using information extracted from an arterialpressure waveform, or from a waveform that is proportional to arterialpressure, which may be measured with an invasive, non-invasive, or“minimally invasive” instrument or combination of instruments Theorganization of this description is: First, the theoretical basis of theinvention is discussed. This is followed by an explanation of the mainsteps of a method to use the theory, then a description of a system thatimplements the method.

The invention may be used to advantage with any type of subject, whetherhuman or animal. Because it is anticipated that the most common use ofthe invention will be on humans in a diagnostic setting, the inventionis described below primarily in use with a “patient.” This is by way ofexample only, however—it is intended that the term “patient” shouldencompass all subjects, both human and animal, regardless of setting.

Theoretical Basis of the Invention

As mentioned above, the principle formula for calculating cardiac output(CO) is CO=SV·HR, where SV is stroke volume and HR is heart rate. GivenHR, the problem then remains how to determine SV. Based on theobservation that the pulsatility of a pressure waveform is created bythe cardiac stroke volume into the arterial tree, inventors havediscovered that one particularly elegant solution to the problem is toestimate SV to be proportional to the standard deviation of the arterialpressure waveform P(t), or of some other signal that itself isproportional to P(t). Thus, according to this aspect of the invention,SV=K·σ _(P)from which follows thatCO=K·σ _(P) ·HRwhere K is a constant and σ_(P) is the standard deviation of P(t) (or ofsome other signal proportional to P(t)) taken over some time interval.Other quickly computed functions of P(t) that provide valuesproportional to SV are also discussed below.Steps for Determining CO

As mentioned above, the invention may be used to determine SV, and henceany other cardiac value that is a function of SV. It is anticipated thatthe most common use of the invention will be to determine CO, so theexemplary embodiments of the invention are described below primarilywith respect to this application. Since the invention generates anestimate of SV, however those skilled in the art of medical systems willknow how to use the estimated SV value to derive still other cardiacvalues.

The first step in the method for measuring CO is to acquire arepresentation of either the arterial pressure waveform P(t) or of somewaveform that is proportional to P(t). This may be a direct measurementof arterial pressure, or a measurement of some other parameter that isproportional to arterial pressure. For example, a plethysmographictechnique, for example, using a finger cuff, will produce a signal thatis proportional to arterial blood pressure; this signal can then bescaled to convert to an estimate of blood pressure. For the sake ofsuccinctness, the method for determining SV is discussed below withreference to the arterial pressure waveform P(t), but one should keep inmind that the disclosed method steps may easily be adjusted to useanother, proportional waveform, simply by applying the proper scaling,such as the scaling provided by a calibration factor K described below.

FIG. 1 illustrates an example of the waveform P(t) of arterial pressuretaken over a single heart cycle, here, from the point of diastolicpressure P_(dia) at time t_(dia0), through the time t_(sys) of systolicpressure P_(sys), to a time t_(dia1) at which the blood pressure onceagain reaches P_(dia).

According to the invention, P(t), or any signal that is proportional toP(t), may be measured at any point in the arterial tree, eitherinvasively or non-invasively. If invasive instruments are used, inparticular, catheter-mounted pressure transducers, then any artery maybe used as a measurement point. Placement of non-invasive transducerswill typically be dictated by the instruments themselves—the placementof finger cuffs, upper arm pressure cuffs, and earlobe clamps should beobvious. Regardless of the instrument, it will ultimately produce, orcause to be produced, an electric signal corresponding (for example,proportional) to P(t).

As is well known, analog signals such as P(t) can be digitized into asequence of digital values using any standard analog-to-digitalconverter (ADC). In other words, P(t), t0≦t≦tf, can be converted, usingknown methods and circuitry, into the digital form P(k), k=0, (n−1),where t0 and tf are initial and final times, respectively, of themeasurement interval and n is the number of samples of P(t) to beincluded in the calculations, distributed usually evenly over themeasurement interval.

The next step of the procedure is to calculate the standard deviation ofP(k), that is, σ_(P), or some value that has a known relationship to thestandard deviation (see below). One way to calculate σ_(P) is to use thewell-known algorithm for calculating the standard deviation of adiscretized function. Thus:

$\begin{matrix}{\sigma_{P}^{2} = {\frac{1}{n - 1}{\sum\limits_{k = 1}^{n}\left\lbrack {{P(k)} - P_{avg}} \right\rbrack^{2}}}} & \left( {{Equation}\mspace{14mu} 1} \right)\end{matrix}$where P_(avg) is the mean pressure value, that is:

$\begin{matrix}{P_{avg} = {\frac{1}{n}{\sum\limits_{k = 1}^{n}{P(k)}}}} & \left( {{Equation}\mspace{14mu} 2} \right)\end{matrix}$Of course, to get σ_(P) the system simply takes the square root of σ_(P)².

The analog measurement interval, that is, the time window [t0, tf], andthus the discrete sampling interval k=1, n, over which a CO estimate iscalculated should be small enough so that it does not encompasssubstantial shifts in the mean pressure P_(avg). Also, one could filterout low frequency variations such as respiration using a high passfilter, which would also help remove the effect of any drift in meanarterial pressure during the time window. For the sake of providing morestable and reliable readings, however, is it best to let the time windowfor each CO estimate extend longer than one cardiac cycle. Preferably,the measurement interval (time window) should be a plurality of cardiaccycles, that is, beginning and ending at the same point in differentcardiac cycles; this ensures that the mean pressure value used in thecalculations of σ will use a mean pressure value P_(avg) that is notskewed because of incomplete measurement of a cycle. Since theinformation is available to any embodiment of the invention that usesP_(avg) (for example, in the calculation of σ_(P)), it may be useful insome implementations to always display the current mean blood pressureand/or trend to the user.

Larger sampling windows have the advantage that the effect ofperturbations such as those caused by reflections will usually bereduced, since they will be tend to “cancel out” in the calculations ofmean pressure and standard deviation. An appropriate time window can bedetermined using normal experimental and clinical methods. Note that itwould be possible for the time window to coincide with a single heartcycle, in which case mean pressure shift will not be of concern.

As a check, the system according to the invention could also, as aseparate background operation, compute the mean pressure over eachcardiac cycle. If the mean cycle-to-cycle pressure shows any absolute orproportional drift greater than some threshold value, a warning signalcould be generated such that the currently computed CO estimate may beconsidered less reliable or discarded altogether.

It would be also possible to adjust the time window [t0, tf] accordingto drift in P_(avg). For example, if P_(avg) over a given time windowdiffers absolutely or proportionately by more than a threshold amountfrom the P_(avg) of the previous time window, then the time window couldbe reduced; stability of P_(avg) could then be used to indicate that thetime window can be expanded. The time window could also be expanded andcontracted based on noise sources, or on a measure of SNR or of SNRvariation. In a preferred embodiment, limits are placed on how much thetime window is allowed to expand or contract; and an indication of thetime interval is displayed to the user.

It is not necessary for the time window to start at any particular pointin the cardiac cycle. Thus, t₀ need not be the same as t_(dia0),although this may be a convenient choice in many implementations. Thismeans that the beginning and end of each measurement interval (that is,t0 and tf), each yielding a CO estimate, may be triggered on almost anycharacteristic of the cardiac cycle, such as at times t_(dia0) ort_(sys), or on non-pressure characteristics such as R waves, etc.

Rather than calculate a single σ_(P) value from a multi-cyclemeasurement as in Equation 1 it would also be possible to calculateseveral σ_(P), for example, one for each of a plurality of intervals,and then to average them (by taking the mean, median, etc.) to compute acomposite σ_(P) for use in the formulas. The inventors have, moreover,discovered other alternatives for computing a pulsatility variablesimilar to the standard deviation σ_(P) of the arterial blood pressure.

The inventors have observed that other values may be derived from thepressure waveform P(t) that either provide an approximation of up, orthat also are proportional to SV, or both. As one example, the inventorshave observed that the difference between the maximum and mini mummeasured pressures, taken over the time window, is a pulsatilitymeasurement that may be substituted for direct calculation of σ_(P)using the standard formulas given above as Equations 1 and 2. Letmax[P(k)] and min[P(k)] be the maximum and minimum values, respectively,of the sampled pressure over the measurement interval. The standarddeviation is approximately equal to one-third times the difference ofthese two values:σ_(P)≈{max[P(k)]−min[P(k)]}/3

Although probably less accurate than calculation of up using Equations 1and 2, this “rough” σ_(P) approximation has the advantage of simplicity,requiring no sampling of P(t) at all. Indeed, given an input signalindicating heart rate (HR), a system to compute {max[P(k)]−min[P(k)]}/3and, from it, SV and/or CO (or some other function of SV) could beimplemented completely in hardware, even all-analog circuitry, usingknown circuit design techniques. This would allow development of veryinexpensive, easily manufactured and physically robust CO monitors foruse in areas or applications that have only minimal facilities andresources.

Of course, it is not necessary to have a separate calculation relatingmax[P(k)] and min[P(k)] to σ_(P), and then to use up to calculate SV.This is described here by way of illustration only. Rather, givenmax[P(k)] and min[P(k)], SV can be estimated directly as:SV=k·{max[P(k)]−min[P(k)]}where k=K/3. (Of course, K can simply be adjusted to account for thefactor ⅓)

As another alternative, the inventors have observed that the maximum orabsolute value of the minimum of the first derivative of the P(t) withrespect to time is generally proportional to σ_(P) and to SV. Thus:

$\begin{matrix}{{SV} = {K \cdot {\max\left( \frac{\mathbb{d}{P(t)}}{\mathbb{d}t} \right)}}} & \; & {or} & \; & {{SV} = {K \cdot {{\min\left( \frac{\mathbb{d}{P(t)}}{\mathbb{d}t} \right)}}}}\end{matrix}$

It would also be possible to use the average or these first derivativesinstead of using only the one or the other. Given P(k), the derivativesmay be determined using any known numerical method; note that the pointsof interest on the pressure waveform are the points of inflection, thatis, the points at which the second time derivative of P(t) is zero. Thetime interval over which these derivatives is evaluated may be theentire cardiac cycle. It will generally suffice, however, to evaluateP(t) between the beginning of the cardiac and the first dicrotic point,shown as P_(dicrotic) in FIG. 1, since the maximum positive slope willusually occur about half way between the diastolic and systolic points,that is, P_(dia) and P_(sys) and the greatest negative slope willgenerally occur about half way between the systolic and first dicroticpoints, that is, P_(sys) and P_(dicrotic). Examining only these portionsof P(t) will eliminate the possibility that spurious values will be usedfrom after the time of P_(dicrotic).

As yet another alternative, a standard software or hardware module couldbe used to compute the Fourier transform of the measured pressure signalP(t) over each cardiac cycle, or over a multiple of cycles. The quantitydefined by the magnitude of the Fourier transform component, H1, at theprimary frequency component, i.e., the frequency corresponding to the“heart rate,” divided by the mean arterial pressure P_(avg), that is,H1/P_(avg), will be proportional to SV. Instead of H1, the magnitude ofthe Fourier component H2 corresponding to twice the heart rate, that is,the first harmonic frequency component, cold be used instead; thus,H2/P_(avg) will also be proportional to SV.

In order to calculate CO, the heart rate HR (or some signal from whichHR can be derived) is needed. Any of the many known instruments formeasuring HR may be used. If the beginning and end times for each P(t)interval are triggered by an electrocardiogram signal, for example, thenthe same signal may be used to calculate HR. The measured pressure waveP(t) (in practice, P(k)) may itself be used to derive HR, for example,using standard Fast Fourier transformation or derivative analysis.

Before finally arriving at a value for CO, it is also necessary todetermine a value for the calibration constant K. One way to do this isas any pre-determined function of P(t); thus, K=K(P(t)). In this case noindependent CO technique is necessary.

Another way to do this is to use any known, independent GO technique todetermine this relationship, whether invasive, for example,thermodilution, or non-invasive, for example, trans-esophagealechocardiography (TEE) or bio-impedance measurement. The inventionprovides continuous trending of CO between intermittent measurementssuch as TD or TEE. Using the chosen independent method, a value CO_(cal)is determined, so that K will be:K=CO _(cal)/(V·HR)where V is the chosen value proportional to SV, for example:

V=σ_(P); or

V=max[P(k)]−min[P(k)]; or

V=maximum or absolute value of the minimum of the first derivative ofthe P(t); or

V=H1/P_(avg) or H2/P_(avg)

Even if an invasive technique such as catheterization is used todetermine K, it will usually not be necessary to leave the catheter inthe patient during the subsequent CO-monitoring session. Moreover, evenwhen using catheter-based calibration technique to determine K, it isnot necessary according to the invention for the measurement to be takenin or near the heart; rather, the calibration measurement could be madein the femoral artery. As such, even where an invasive technique is usedto determine the calibration constant K, the invention as a whole isstill minimally invasive in that any catheterization may be peripheraland temporary.

As is mentioned above, rather than measure arterial blood pressuredirectly, any other input signal may be used that is proportional toblood pressure. This means that calibration may be done at any or all ofseveral points in the calculations. For example, if some signal otherthan arterial blood pressure itself is used as input, then it may becalibrated to blood pressure before its values are used to calculatestandard deviation, or afterwards, in which case either the resultingstandard deviation value can be scaled, or the resulting SV value can becalibrated (for example, by setting K properly), or some final functionof SV (such as CO) can be scaled. In short, the fact that the inventionmay in some cases use a different input signal than a direct measurementof arterial blood pressure does not limit its ability to generate anaccurate SV estimate

System Components

FIG. 2 shows the main components of a system that implements the methoddescribed above for sensing pressure and calculating CO. As is mentionedabove, pressure, or some other input signal proportional to pressure,may be sensed in either or, indeed, both, of two ways: invasively andnon-invasively. Simply because it is anticipated to be the most commonimplementation of the invention, the system is described as measuringarterial blood pressure (as opposed to some other input signal that isconverted to pressure) and generating a CO estimate (as opposed to justSV). Changes to the illustrated system to accommodate other designchoices will be obvious to those skilled in the art of medical devices.

FIG. 2 shows both types of pressure sensing for the sake of conciseness;in most practical applications of the invention, either one or severalvariations will typically be implemented. In invasive applications ofthe invention, a conventional pressure sensor 100 is mounted on acatheter 110, which is inserted in an artery 120 of a portion 130 of thebody of a human or animal patient. Such artery could be an ascendingaorta, or pulmonary artery, or, in order to reduce the level ofinvasiveness, the artery 120 could be peripheral, such as the femoral,radial or brachial artery. In the non-invasive applications of theinvention, a conventional pressure sensor 200, such as aphoto-plethysmographic blood pressure probe, is mounted externally inany conventional manner, for example using a cuff around a finger 230 ora transducer mounted on the wrist of the patient. FIG. 2 schematicallyshows both types.

The signals from the sensors 100, 200 are passed via any knownconnectors as inputs to a processing system 300, which includes one ormore processors and other supporting hardware and system software (notshown) usually included to process signals and execute code. Theinvention may be implemented using a modified, standard, personalcomputer, or it may be incorporated into a larger, specializedmonitoring system. In this invention, the processing system 300 also mayinclude, or is connected to, conditioning circuitry 302 which performssuch normal signal processing tasks as amplification, filtering,ranging, etc., as needed, as well as the optional high pass filteringmentioned above. The conditioned, sensed input pressure signal P(t) isthen converted to digital form by a conventional analog-to-digitalconverter ADC 304. As is well understood, the sampling frequency of theADC 304 should be chosen with regard to the Nyquist criterion so as toavoid aliasing of the pressure signal; this procedure is very well knownin the art of digital signal processing. The output from the ADC 304will be the discrete pressure signal P(k), whose values may be stored inconventional memory circuitry (not shown).

The values P(k) are passed to (usually, accessed from memory by) to anSV-calculation module 306, which is a software component comprisingprocessor-executable code for calculating whichever value V is used todetermine SV as explained above. For example, where up is calculateddirectly, the SV-calculation module 306 will evaluate Equations 1 and 2above, or equivalent expressions. The calculation module 306 preferablyalso selects the time window [t0, tf] over which each CO estimate isgenerated. This may be done as simply as choosing which and how many ofthe stored, consecutive, discretized P(t) values P(k) are used in eachcalculation, which is the same as selecting n in the range k=1, . . . ,n.

The computer value of SV is then passed to a subsequent CO-calculationmodule 308, which similarly comprises executable code for evaluate theexpression CO=K·σ_(P)·HR. Of course, modules 306 and 308 may be combinedinto a single software component; they are shown separately for the sakeof clarity. Of course, the CO-calculation module 308 requires values forHR and K as well:

The patient's current heart rate HR is either calculated from themeasured pressure curve P(k) by a corresponding software module 310 (forexample, using Fourier or derivative analysis) or is otherwise measuredwith any conventional hardware device.

The value K will normally be input to module 308 automatically, but canbe entered by the operator via a conventional input device 400 such as akeyboard, mouse, etc. This will usually be the same input device(s) usedby the processing system 300 for other purposes such as entering dataidentifying the patient and the specifications of the monitoringsession.

The module 308 calculates and estimates CO for each chosen measurementinterval. Each estimated CO value is preferably output to anyconventional display or printing device 500 for the user to view andmonitor. As with the input device 400, the display 500 will typically bethe same as is used by the processing system for other purposes.

The invention further relates to a computer program loadable in acomputer unit or the processing system 300 in order to execute themethod of the invention. Moreover, the various software modules 306,308, 310, and 312 used to perform the various calculations and performrelated method steps according to the invention may also be stored ascomputer-executable instructions on a computer-readable medium in orderto allow the invention to be loaded into and executed by differentprocessing systems.

Other Outputs

The invention is described above in the context of calculating estimatesof CO. This is the use of invention that the inventor assumes will bemost common, but the invention is not inherently limited to such use. Inessence, the invention provides a novel way to calculate stroke volumeSV, or any parameter that is a function of (for example, proportionalto) SV, not just CO. Consequently, the advantages of the invention willapply to the calculation of any value derived from SV. For example, theend diastolic volume (EDV) and the ejection fraction (EF) are related asEF=SV/EDV, which expresses the intuitive relationship that the pumpingefficiency (EF) of the heart is the ratio between how much blood theheart pumps out on every beat (contraction) and how much blood is in theheart chamber just before the beat. Inversely, EDV=SV/EF. If an estimatefor either EDV or EF is determined in some other known manner, then theinvention could be used to provide SV, and thus an estimate of the otherof EDV or EF.

1. A method for determining cardiac stroke volume of a subjectcomprising: sensing arterial blood pressure; converting the sensedarterial blood pressure to a pressure signal; determining the heart rateof the subject; determining the amplitude of a spectral component of thepressure signal for a frequency corresponding to a multiple of the heartrate; calculating an average value of the pressure signal; andcalculating an estimate of the stroke volume as a function of the ratioof the amplitude of the spectral component and the average value of thepressure signal.
 2. A method of claim 1, wherein sensing uses a bloodpressure sensor, converting uses an analog-to-digital converter andfurther comprising outputting a parameter based on the estimate ofstroke volume on a display.
 3. A method of claim 1, wherein determiningthe amplitude of the spectral component includes using a Fouriertransform.
 4. A method of claim 3, wherein determining the heart rateincludes using a Fourier transform.
 5. A method of claim 4, whereinsensing arterial pressure is performed using a sensor placed in aperipheral artery.
 6. A method for determining cardiac stroke volume ofa subject comprising: sensing arterial blood pressure; converting thesensed arterial blood pressure to a pressure signal; detecting a maximumvalue and minimum value of the first time derivative of the pressurevalue during a measurement interval; and calculating an estimate of thestroke volume as a function of a magnitude of one of the maximum orminimum values, wherein the estimate of the stroke volume isproportionate to a standard deviation calculated using the magnitude ofthe maximum and minimum values.
 7. A method of claim 6, wherein theestimate of the stroke volume is also proportionate to the magnitude ofthe maximum or minimum value.
 8. A method of claim 7, wherein theestimate of the stroke volume is proportionate to the magnitude of boththe maximum and minimum values.
 9. A method of claim 6, wherein thestandard deviation is calculated by dividing a sum of the magnitude ofthe maximum and minimum values by
 3. 10. A method of claim 6, whereindetection of the maximum and minimum values is determined only betweenthe beginning of a cardiac cycle and a dichrotic point.
 11. A method ofclaim 6, wherein sensing uses a blood pressure sensor, converting usesan analog-to-digital converter and further comprising outputting aparameter based on the estimate of stroke volume on a display.
 12. Amethod of claim 6, wherein a component of the estimate of the strokevolume includes the magnitude of one of the maximum or minimum values.